Magnetic resonance imaging (MRI) involves the transmission and receipt of radio frequency (RF) energy. RF energy may be transmitted by an RF coil. Resulting nuclear magnetic resonance (NMR) signals may also be received by an RF coil. In early MRI, RF energy may have been transmitted from a single RF coil and resulting NMR signals received by a single RF coil. Later, multiple receivers may have been used in parallel acquisition techniques. Similarly, multiple transmitters may have been used in parallel transmission techniques.
RF coils create the B1 field that rotates the net magnetization in a pulse sequence. RF coils may also detect precessing transverse magnetization. Thus, RF coils may be transmit coils, receive coils, or transmit and receive (transmit/receive) coils. An RF coil may include, for example, an LC circuit. Typically, the transmitted RF signals are orders of magnitude larger than the MR signals generated by the excited nuclei and detected by the RF receive coils. Modern MRI apparatus may include multiple RF coils arranged in a receive array. An RF coil receive array may include hundreds or thousands of delicate electronic components. Damage to just one electronic component in the RF coil may cause the RF coil to no longer function in its optimized and expected way. To protect the receive coils and receiver circuits and apparatus, the receive coils may be decoupled or detuned while RF is being transmitted by an MR apparatus. Decoupling an RF coil reduces the current flowing through the RF coil. The decoupling or detuning may be active or passive. One approach to active decoupling involves, for example, applying a bias to a PIN diode semiconductor switch in conjunction with an LC circuit during RF transmission. Passive decoupling involves, for example, using antiparallel diode semiconductor switches in conjunction with LC circuitry. The antiparallel diode semiconductor switches are switched upon detecting high power RF transmit pulses, which allows high induced voltage generated from transmitting fields, but not low strength signals from nuclei, to interact with the parallel resonant LC circuit that decouples the coil.
An imaging coil needs to be able to resonate at a selected Larmor frequency. Imaging coils include inductive elements and capacitive elements. The resonant frequency, v, of an RF coil is determined by the inductance (L) and capacitance (C) of the inductor capacitor circuit (e.g. LC circuit) according to:
  v  =      1          2      ⁢                          ⁢      Π      ⁢              LC            
Imaging coils may need to be tuned. Tuning an imaging coil may include varying the performance of a capacitor. Recall that frequency: f=ω/(2π), wavelength, λ=c/f, and λ=4.7 m at 1.5 T. Recall also that the Larmor frequency: f0=γ B0/(2π), where γ/(2π)=42.58 MHz/T; at 1.5 T, f0=63.87 MHz; at 3 T, f0=127.73 MHz; at 7 T, f0=298.06 MHz. Basic circuit design principles include the fact that capacitors add in parallel (impedance 1/(jCω)) and inductors add in series (impedance jLω).
When MRI coils that are tuned to the same radio frequency are positioned close together, which may occur, for example, in phased array coils, the MRI coils may inductively couple to each other, which causes the MRI coils to detune each other. Detuning due to inductive coupling reduces image quality as compared to using single coils individually. Conventional phased array coils may address the detuning due to inductive coupling problem by overlapping coils or by using preamplifiers that dampen current flow in individual coils.
FIG. 1 illustrates a single RF coil segment 102 shown schematically to include an inductance 103, a resistance 104, and a capacitance 105. Capacitance 105 is selected to tune the segment 102 to a desired frequency (e.g., Larmor frequency). The RF coil segment 102 is connected across the output of a current control circuit 106 that is driven by an input signal 107 to produce a current in the RF coil segment 102. Unfortunately, an additional induced current may also flow through the RF coil segment 102 due to signals indicated at 108 induced by currents flowing in other (e.g., adjacent) RF coil segments. With multiple driving loops tuned at a single image frequency, which may occur in a phased array coil, the current on a loop is a superposition of the driven current and currents induced by other transmitters due to electromagnetic induction.
Conventional approaches to reduce, minimize, or eliminate the coupling through the mutual impedance in two interacting elements may have been attempted by cancelling the mutual impedance or by reducing the current in the coil. Mutual inductance may be cancelled by either partial overlap of adjacent coils, which may be referred to as transformer type decoupling, or by using decoupling capacitors. Other conventional approaches include using a preamplifier decoupling network to isolate coil elements by creating a large impedance block at the terminals of a receive element, which suppresses currents driven by the spin induced electromotive force (emf).
There are many design issues associated with MRI RF coil design. For example, the inductance of a conventional coil depends on the geometry of the coil. For a square coil with a side length a and wire diameter f: L=[μ0/π] [−4a+2a √2−2a log(1+√2)+2a log(4a/f)]. For a loop coil with loop diameter d and wire diameter f: L=[μ0d/2] [log(8d/f)−2]. Thus, the selection of the geometry of a coil determines, at least in part, the inductance of the coil.
The resistance of a coil also depends on the geometry of the coil. The resistance R of a conductor of length l and cross-sectional area A is R=ρl/A, where ρ is the conductor resistivity and is a property of the conductor material and the temperature. For a copper wire coil with loop diameter d and wire diameter f: R=dρCu/(fδCu). For a copper foil coil with loop diameter d, copper thickness t, and copper width w: R=πdρCu/(2wδCu), where t is much greater than the copper skin depth and w is much greater than t. Thus, the selection of the geometry of a coil and the material (e.g., wire, foil) determines, at least in part, the inductance of the coil. The length of the loop also impacts the properties of the coil.
Coils may be characterized by their signal voltage, which is the electro-motive force (emf) induced in a coil: ξ=−∂φ/∂t∝−∂(B1·M0)/∂t, where φ is the magnetic flux across the coil (closed loop), magnetization M0=Nγ2(h/(2π))2s(s+1)B0/(3kBTS)=σ0B0/μ0, where N is the number of nuclear spins s per unit volume (s=½ for protons) and Ts is the temperature of the sample. Since ω0=γB0, ξ∝ω02. The noise in a coil may be thermal (e.g., v=(4kBTSRΔf)1/2, where R is the total resistance, kB is the Boltzmann constant, T is the temperature in K, and Δf is the bandwidth of the received signal). The signal to noise ratio (SNR) for a coil may be described by ξ/v.
Coils may be used for transmitting RF energy that is intended to cause nuclear magnetic resonance (NMR) in a sample. The frequency at which NMR will be created depends on the magnetic field present in the sample. Both the main magnetic field B0 produced by the MRI apparatus and the additional magnetic field B1 produced by a coil contribute to the magnetic field present in the sample. For a circular loop coil, the transmit B1 field equals the coil sensitivity. A circular loop of radius a carrying a current I produces on axis the field: B=μ0 I a2/[2(a2+z2)3/2].
RF coils for MRI may need to be tuned and matched. Tuning involves establishing or manipulating the capacitance in a coil so that a desired resistance is produced. Matching involves establishing or manipulating the capacitance in a coil so that a desired reactance is achieved. When tuning, the impedance z may be described by Z=R+jX=1/(1/(r+jLω)+jCω). Tuning may be performed to achieve a desired tuning frequency for a coil. ω0 identifies the desired tuning frequency. ω0, may be, for example, 63.87 MHz at 1.5 Tesla (T). The size of a conventional coil facilitates estimating inductance L. With an estimate of L in hand, values for capacitors can be computed to produce a desired resonant peak in an appropriate location with respect to ω0. Once capacitors are selected, the resonant peak can be observed and a more accurate L can be computed. The capacitors can then be adjusted to produce the desired resistance. Once the desired resistance is achieved, then capacitance can be adjusted to cancel reactance.
Conventional coils may use PIN diodes. When forward-biased, a PIN diode may produce a negligible resistance (e.g., 0.5Ω), which is effectively a short-circuit. When reverse-biased, a PIN diode may produce a high resistance (e.g., 200 kΩ) in parallel with a low capacitance (e.g., ˜2 pF), which is essentially an open-circuit. RF coils may employ PIN diodes as part of detuning circuits that facilitate disabling a receive function of the RF coil while the RF coil is transmitting.
FIG. 2 illustrates a schematic of a simple conventional RF coil 200 for MRI. The coil 200 is illustrated as a loop 210. Loop 210 has elements that produce a resistance (R) (e.g., resistor 220) and that produce an inductance (L) (e.g., inductor 230). A conventional loop may include a matching capacitor 240 and tuning capacitor 250 that produce capacitance (C). The simple RF coil 200 may be referred to as an LC coil or as an RLC coil. Conventionally, the resistor 220, inductor 230, and capacitor 250 may all have been two-terminal passive elements that were soldered to copper wire or copper foil that was attached to a printed circuit board.
A resistor may be, for example, a passive, two-terminal electrical component that implements electrical resistance as a circuit element. Resistors reduce current flow. Resistors also lower voltage levels within circuits. Resistors may have fixed resistances or variable resistances. The current that flows through a resistor is directly proportional to the voltage applied across the resistor's terminals. This relationship is represented by Ohm's Law: V=IR, where I is the current through the conductor, V is the potential difference across the conductor, and R is the resistance of the conductor.
An inductor may be a passive two-terminal electrical component that resists changes in electric current. An inductor may be made from, for example, a wire that is wound into a coil. When a current flows through the inductor, energy may be stored temporarily in a magnetic field in the coil. When the current flowing through the inductor changes, the time-varying magnetic field induces a voltage in the inductor. The voltage will be induced according to Faraday's law and thus may oppose the change in current that created the voltage.
A capacitor may be a passive, two-terminal electrical component that is used to store energy. The energy may be stored electrostatically in an electric field. Although there are many types of capacitors, capacitors tend to contain at least two electrical conductors that are separated by a dielectric material. The conductors may be, for example, plates and the dielectric material may be, for example, an insulator. The conductors may be, for example, thin films of metal, aluminum foil or other materials. The non-conducting dielectric material increases the capacitor's charge capacity. The dielectric material may be, for example, glass, ceramic, plastic film, air, paper, mica, or other materials. Unlike a resistor, a capacitor does not dissipate energy. Instead, a capacitor stores energy in the form of an electrostatic field between its conductors.
When there is a potential difference across the conductors, an electric field may develop across the dielectric material. The electric field may cause positive charge (+Q) to collect on one conductor and negative charge (−Q) to collect on the other conductor.
FIG. 3 illustrates a schematic of another simple RF coil 300 for MRI. RF coil 300 may also be referred to as an LC coil or as an RLC coil. The RF coil 300 is illustrated as a square loop 310. Loop 310 has elements that produce a resistance (e.g., resistor 320) and elements that produce an inductance (e.g., inductor 330). A conventional loop may include a capacitor 340 and capacitor 350 that work together to achieve matching. The resistor 320, inductor 330, and capacitors 340 and 350 may have been soldered to copper wire or copper foil that was attached to a printed circuit board. RF coil 300 is contrasted with RF coil 200 (FIG. 2) that included capacitor 250 for tuning purposes.
FIG. 4 illustrates a conventional RF coil 400 that performs traditional “distributed” decoupling using components L1 and D1. RF coil 400 includes capacitors C1, C2, C3, and C4, and inductors L1 and L2. RF coil 400 includes a pre-amplifier circuit 410. Conventionally, pre-amplifier circuit 410 may be used as a pre-amplifier decoupling circuit. When the input impedance of the pre-amplifier circuit 410 becomes small (e.g. 0 to 2 Ohms) inductor L2 achieves a parallel resonant circuit with C3. The pre-amplifier, in conjunction with a coil matching/decoupling circuit, reduces the current flow in the RF coil. Reducing the current flow in the RF coil reduces the magnetic fields generated by the current flow, which would otherwise induce undesirable current flow in the neighboring or adjacent RF coils.
RF coil 400 also includes a PIN diode D1. Recall that a PIN diode has a wide, lightly doped near intrinsic semiconductor region positioned between a p-type semiconductor region and an n-type semiconductor region that are used for ohmic contacts. The wide intrinsic region makes the PIN diode suitable for fast switches. Fast switching may be employed in MRI coils. In transmit mode, the PIN diode D1 may be turned on (e.g., shorted).
In conventional RF coil 400, capacitors C1, C2, and C3 are illustrated to represent one or more capacitors that may be employed in the RF coil 400. Thus, capacitors C1-C3 may be an equivalent capacitor of multiple breaking point capacitors that may appear in RF coil 400 minus capacitor C4. The RF coil has an inductance. The inductance may be produced, for example, by a copper trace that forms the RF coil 400.
In the conventional RF coil 400, capacitor C4 may be the breaking point capacitor that is used for decoupling the RF coil 400 from other MRI coils. Capacitor C4 and inductor L1 are in parallel resonance and the impedance across capacitor C4 is high. Capacitor C4 is a single high impedance point in RF coil 400. Since the impedance across capacitor C4 is high, an induced voltage on RF coil 400 cannot generate a large current through capacitors C1, C2, or C3.